The present invention relates to improved apparatus and methods for canceling feedback in audio systems such as hearing aids.
Mechanical and acoustic feedback limits the maximum gain that can be achieved in most hearing aids (Lybarger, S. F., xe2x80x9cAcoustic feedback controlxe2x80x9d, The Vanderbilt Hearing-Aid Report, Studebaker and Bess, Eds., Upper Darby, Pa.: Monographs in Contemporary Audiology, pp 87-90, 1982). System instability caused by feedback is sometimes audible as a continuous high-frequency tone or whistle emanating from the hearing aid. Mechanical vibrations from the receiver in a high-power hearing aid can be reduced by combining the outputs of two receivers mounted back-to-back so as to cancel the net mechanical moment; as much as 10 dB additional gain can be achieved before the onset of oscillation when this is done. But in most instruments, venting the BTE earmold or ITE shell establishes an acoustic feedback path that limits the maximum possible gain to less than 40 dB for a small vent and even less for large vents (Kates, J. M., xe2x80x9cA computer simulation of hearing aid response and the effects of ear canal sizexe2x80x9d, J. Acoust. Soc. Am., Vol. 83, pp 1952-1963, 1988). The acoustic feedback path includes the effects of the hearing-aid amplifier, receiver, and microphone as well as the vent acoustics.
The traditional procedure for increasing the stability of a hearing aid is to reduce the gain at high frequencies (Ammitzboll, K., xe2x80x9cResonant peak controlxe2x80x9d, U.S. Pat. No. 4,689,818, 1987). Controlling feedback by modifying the system frequency response, however, means that the desired high-frequency response of the instrument must be sacrificed in order to maintain stability. Phase shifters and notch filters have also been tried (Egolf, D. P., xe2x80x9cReview of the acoustic feedback literature from a control theory point of viewxe2x80x9d, The Vanderbilt Hearing-Aid Report, Studebaker and Bess, Eds., Upper Darby, Pa.: Monographs in Contemporary Audiology, pp 94-103, 1982), but have not proven to be very effective.
A more effective technique is feedback cancellation, in which the feedback signal is estimated and subtracted from the microphone signal. Computer simulations and prototype digital systems indicate that increases in gain of between 6 and 17 dB can be achieved in an adaptive system before the onset of oscillation, and no loss of high-frequency response is observed (Bustamante, D. K., Worrell, T. L., and Williamson, M. J., xe2x80x9cMeasurement of adaptive suppression of acoustic feedback in hearing aidsxe2x80x9d, Proc. 1989 Int. Conf. Acoust. Speech and Sig. Proc., Glasgow, pp 2017-2020, 1989; Engebretson, A. M., O""Connell, M. P., and Gong, F., xe2x80x9cAn adaptive feedback equalization algorithm for the CID digital hearing aidxe2x80x9d, Proc. 12th Annual Int. Conf. of the IEEE Eng. in Medicine and Biology Soc., Part 5, Philadelphia, Pa., pp 2286-2287, 1990; Kates, J. M., xe2x80x9cFeedback cancellation in hearing aids: Results from a computer simulationxe2x80x9d, IEEE Trans. Sig. Proc., Vol.39, pp 553-562, 1991; Dyrlund, O., and Bisgaard, N., xe2x80x9cAcoustic feedback margin improvements in hearing instruments using a prototype DFS (digital feedback suppression) systemxe2x80x9d, Scand. Audiol., Vol. 20, pp 49-53, 1991; Engebretson, A. M., and French-St. George, M., xe2x80x9cProperties of an adaptive feedback equalization algorithmxe2x80x9d, J. Rehab. Res. and Devel., Vol. 30, pp 8-16, 1993; Engebretson, A. M., O""Connell, M. P., and Zheng, B., xe2x80x9cElectronic filters, hearing aids, and methodsxe2x80x9d, U.S. Pat. No. 5,016,280; Williamson, M. J., and Bustamante, D. K., xe2x80x9cFeedback suppression in digital signal processing hearing aids,xe2x80x9d U.S. Pat. No. 5,019,952).
In laboratory tests of a wearable digital hearing aid (French-St. George, M., Wood, D. J., and Engebretson, A. M., xe2x80x9cBehavioral assessment of adaptive feedback cancellation in a digital hearing aidxe2x80x9d, J. Rehab. Res. and Devel., Vol. 30, pp 17-25, 1993), a group of hearing-impaired subjects used an additional 4 dB of gain when adaptive feedback cancellation was engaged and showed significantly better speech recognition in quiet and in a background of speech babble. Field trials of a feedback-cancellation system built into a BTE hearing aid have shown increases of 8-10 dB in the gain used by severely-impaired subjects (Bisgaard, N., xe2x80x9cDigital feedback suppression: Clinical experiences with profoundly hearing impairedxe2x80x9d, In Recent Developments in Hearing Instrument Technology: 15th Danavox Symposium, Ed. by J. Beilin and G. R. Jensen, Kolding, Denmark, pp 370-384, 1993) and increases of 10-13 dB in the gain margin measured in real ears (Dyrlund, O., Henningsen, L. B., Bisgaard, N., and Jensen, J. H., xe2x80x9cDigital feedback suppression (DFS): Characterization of feedback-margin improvements in a DFS hearing instrumentxe2x80x9d, Scand. Audiol., Vol. 23, pp 135-138, 1994).
In some systems, the characteristics of the feedback path are estimated using a noise sequence continuously injected at a low level (Engebretson and French-St.George, 1993; Bisgaard, 1993, referenced above). The weight update of the adaptive filter also proceeds on a continuous basis, generally using the LMS algorithm (Widrow, B., McCool, J. M., Larimore, M. G., and Johnson, C. R., Jr., xe2x80x9cStationary and nonstationary learning characteristics of the LMS adaptive filterxe2x80x9d, Proc. IEEE, Vol. 64, pp 1151-1162, 1976). This approach results in a reduced SNR for the user due to the presence of the injected probe noise. In addition, the ability of the system to cancel the feedback may be reduced due to the presence of speech or ambient noise at the microphone input (Kates, 1991, referenced above; Maxwell, J. A., and Zurek, P. M., xe2x80x9cReducing acoustic feedback in hearing aidsxe2x80x9d, IEEE Trans. Speech and Audio Proc., Vol. 3, pp 304-313, 1995). Better estimation of the feedback path will occur if the hearing-aid processing is turned off during the adaptation so that the instrument is operating in an open-loop rather than closed-loop mode while adaptation occurs (Kates, 1991). Furthermore, for a short noise burst used as the probe in an open-loop system, solving the Wiener-Hopf equation (Makhoul, J. xe2x80x9cLinear prediction: A tutorial review,xe2x80x9d Proc. IEEE, Vol. 63, pp 561-580, 1975) for the optimum filter weights can result in greater feedback cancellation than found for LMS adaptation (Kates, 1991). For stationary conditions up to 7 dB of additional feedback cancellation is observed solving the Wiener-Hopf equation as compared to a continuously-adapting system, but this approach can have difficulty in tracking a changing acoustic environment because the weights are adapted only when a decision algorithm ascertains the need and the bursts of injected noise can be annoying (Maxwell and Zurek, 1995, referenced above).
A simpler approach is to use a fixed approximation to the feedback path instead of an adaptive filter. Levitt, H., Dugot, R. S., and Kopper, K. W., xe2x80x9cProgrammable digital hearing aid systemxe2x80x9d, U.S. Pat. No. 4,731,850, 1988, proposed setting the feedback cancellation filter response when the hearing aid was fitted to the user. Woodruff, B. D., and Preves, D. A., xe2x80x9cFixed filter implementation of feedback cancellation for in-the-ear hearing aidsxe2x80x9d, Proc. 1995 IEEE ASSP Workshop on Applications of Signal Processing to Audio and Acoustics, New Paltz, N.Y., paper 1.5, 1995, found that a feedback cancellation filter constructed from the average of the responses of 13 ears gave an improvement of 6-8 dB in maximum stable gain for an ITE instrument, while the optimum filter for each ear gave 9-11 dB improvement.
A need remains in the art for apparatus and methods to eliminate xe2x80x9cwhistlingxe2x80x9d due to feedback in unstable hearing-aids.
The primary objective of the feedback cancellation processing of the present invention is to eliminate xe2x80x9cwhistlingxe2x80x9d due to feedback in an unstable hearing-aid amplification system. The processing should provide an additional 10 dB of allowable gain in comparison with a system not having feedback cancellation. The presence of feedback cancellation should not introduce any artifacts in the hearing-aid output, and it should not require any special understanding on the part of the user to operate the system.
The feedback cancellation of the present invention uses a cascade of two adaptive filters along with a short bulk delay. The first filter is adapted when the hearing aid is turned on in the ear. This filter adapts quickly using a white noise probe signal, and then the filter coefficients are frozen. The first filter models those parts of the hearing-aid feedback path that are assumed to be essentially constant while the hearing aid is in use, such as the microphone, amplifier, and receiver resonances, and the basic acoustic feedback path.
The second filter adapts while the hearing aid is in use and does not use a separate probe signal. This filter provides a rapid correction to the feedback path model when the hearing aid goes unstable, and more slowly tracks perturbations in the feedback path that occur in daily use such as caused by chewing, sneezing, or using a telephone handset. The bulk delay shifts the filter response so as to make the most effective use of the limited number of filter coefficients.
A hearing aid according to the present comprises a microphone for converting sound into an audio signal, feedback cancellation means including means for estimating a physical feedback signal of the hearing aid, and means for modeling a signal processing feedback signal to compensate for the estimated physical feedback signal, subtracting means, connected to the output of the microphone and the output of the feedback cancellation means, for subtracting the signal processing feedback signal from the audio signal to form a compensated audio signal, a hearing aid processor, connected to the output of the subtracting means, for processing the compensated audio signal, and a speaker, connected to the output of the hearing aid processor, for converting the processed compensated audio signal into a sound signal.
The feedback cancellation means forms a feedback path from the output of the hearing aid processing means to the input of the subtracting means and includes a first filter for modeling near constant factors in the physical feedback path, and a second, quickly varying, filter for modeling variable factors in the feedback path. The first filter varies substantially slower than the second filter.
In a first embodiment, the first filter is designed when the hearing aid is turned on and the design is then frozen. The second filter is also designed when the hearing aid is turned on, and adapted thereafter based upon the output of the subtracting means and based upon the output of the hearing aid processor.
The first filter may be the denominator of an IIR filter and the second filter may be the numerator of said IIR filter. In this case, the first filter is connected to the output of the hearing aid processor, for filtering the output of the hearing aid processor, and the output of the first filter is connected to the input of the second filter, for providing the filtered output of the hearing aid processor to the second filter.
Or, the first filter might be an IIR filter and the second filter an FIR filter.
The means for designing the first filter and the means for designing the second filter comprise means for disabling the input to the speaker means from the hearing aid processing means, a probe for providing a test signal to the input of the speaker means and to the second filter, means for connecting the output of the microphone to the input of the first filter, means for connecting the output of the first filter and the output of the second filter to the subtraction means, means for designing the second filter based upon the test signal and the output of the subtraction means, and means for designing the first filter based upon the output of the microphone and the output of the subtraction means.
The means for designing the first filter may further include means for detuning the filter, and the means for designing the second filter may further include means for adapting the second filter to the detuned first filter.
In a second embodiment, the hearing aid includes means for designing the first filter when the hearing aid is turned on, means for designing the second filter when the hearing aid is turned on, means for slowly adapting the first filter, and means for rapidly adapting the second filter based upon the output of the subtracting means and based upon the output of the hearing aid processing means.
In the second embodiment, the means for adapting the first filter might adapts the first filter based upon the output of the subtracting means, or based upon the output of the hearing aid processing means.
A dual microphone embodiment of the present invention hearing aid comprises a first microphone for converting sound into a first audio signal, a second microphone for converting sound into a second audio signal, feedback cancellation means including means for estimating physical feedback signals to each microphone of the hearing aid, and means for modeling a first signal processing feedback signal to compensate for the estimated physical feedback signal to the first microphone and a second signal processing feedback signal to compensate for the estimated physical feedback signal to the second microphone, means for subtracting the first signal processing feedback signal from the first audio signal to form a first compensated audio signal, means for subtracting the second signal processing feedback signal from the second audio signal to form a second compensated audio signal, beamforming means, connected to each subtracting means, to combine the compensated audio signals into a beamformed signal, a hearing aid processor, connected to the beamforming means, for processing the beamformed signal, and a speaker, connected to the output of the hearing aid processing means, for converting the processed beamformed signal into a sound signal.
The feedback cancellation means includes a slower varying filter, connected to the output of the hearing aid processing means, for modeling near constant environmental factors in one of the physical feedback paths, a first quickly varying filter, connected to the output of the slower varying filter and providing an input to the first subtraction means, for modeling variable factors in the first feedback path, and a second quickly varying filter, connected to the output of the slowly varying filter and providing an input to the second subtraction means, for modeling variable factors in the second feedback path. The slower varying filter varies substantially slower than said quickly varying filters.
In a first version of the dual microphone embodiment, the hearing aid further includes means for designing the slower varying filter when the hearing aid is turned on, and means for freezing the slower varying filter design. It also includes means for designing the first and second quickly varying filters when the hearing aid is turned on, means for adapting the first quickly varying filter based upon the output of the first subtracting means and based upon the output of the hearing aid processing means, and means for adapting the second quickly varying filter based upon the output of the second subtracting means and based upon the output of the hearing aid processing means.
In this embodiment, the first quickly varying filter might be the denominator of a first IIR filter, the second quickly varying filter might be the denominator of a second IIR filter, and the slower varying filter might be based upon the numerator of at least one of these IIR filters. Or, the slower varying filter might be an IIR filter and the rapidly varying filters might be FIR filters.
In the dual microphone embodiment, the means for designing the slower varying filter and the means for designing the rapidly varying filters might comprise means for disabling the input to the speaker means from the hearing aid processing means, probe means for providing a test signal to the input of the speaker means and to the rapidly varying filters, means for connecting the output of the first microphone to the input of the slower varying filter, means for connecting the output of the slower varying filter and the output of the first rapidly varying filter to the first subtraction means, means for designing the first rapidly varying filter based upon the test signal and the output of the first subtraction means, means for connecting the output of the slower varying filter and the output of the second rapidly varying filter to the second subtraction means, means for designing the second rapidly varying filter based upon the test signal and the output of the second subtraction means, and means for designing the slower varying filter based upon the output of the microphone and the output of at least one of the subtraction means.
The means for designing the slower varying filter might further include means for detuning the slower varying filter, and the means for designing the quickly varying filters might further include means for adapting the quickly varying filters to the detuned slower varying filter.
Another version of the dual microphone embodiment might include means for designing the slower varying filter when the hearing aid is turned on, means for designing the quickly varying filters when the hearing aid is turned on, means for slowly adapting the slower varying filter, means for rapidly adapting the first quickly varying filter based upon the output of the first subtracting means and based upon the output of the hearing aid processing means, and means for rapidly adapting the second quickly varying filter based upon the output of the second subtracting means and based upon the output of the hearing aid processing means.
In this case, the means for adapting the slower varying filter might adapt the slower varying filter based upon the output of at least one of the subtracting means, or might adapt the slower varying filter based upon the output of the hearing aid processing means.
Improvements to the feedback cancellation processing of the present invention include improvements to the fitting and initialization of the hearing aid, and improvements to the feedback cancellation processing. With regard to fitting and initializing the feedback cancellation hearing aid, the feedback path model determined during initialization may be used to set the maximum gain allowable in the hearing aid. This maximum stable gain can be used to assess the validity of the hearing aid design, by determining whether the recommended gain for that design exceeds the maximum stable gain. Further, the hearing aid fitting in the ear canal may be tested for leakage, by testing whether the maximum stable gain computed for the hearing aid with its vent hole blocked is substantially higher than the maximum stable gain computed for the hearing aid with its vent open.
Another fitting and initialization feature allows the use of the error signal plotted versus time in the feedback cancellation system as a convergence check of the system, or the amount of feedback cancellation can be estimated by comparing the error at the end of convergence to that at the start of convergence. The error signal may also be used to do an iterative selection of optimum bulk delay in the feedback path, with the optimum delay being that which gives the minimum convergence error. Or, the bulk delay may be set by choosing a preliminary delay, allowing the zero model coefficients to adapt, and adjusting the preliminary delay so that the coefficient having the largest magnitude is positioned at a desired tap location.
With regard to the feedback cancellation processing, the amplitude of the noise probe signal may be adjusted in response to the ambient noise level in the room (this could also be done as part of initialization and fitting). Another processing improvement involves adding a 0 Hz blocking filter as a fixed component to the feedback path, to remove DC bias. In another improvement, the hearing aid gain may be adjusted as a function of the zero coefficient vector.
Another feedback cancellation processing feature allows the LMS adaptation step size to be adjusted in response to an estimate of the input power to the hearing aid. This power estimate may also be used to determine whether the LMS zero filter update is likely to overflow the accumulator. As another feature, the output power is tested to determine whether distortion is likely.
Another feedback cancellation processing feature replaces the adaptive zero filter with an adaptive gain. In another improvement, the pole filter may be improved by switching or interpolating between two sets of frozen filter coefficients. Another processing feature constrains the gain of the adaptive feedback path filter.